MR imaging of internal body tissues may be used for numerous medical procedures, including diagnosis and surgery. In general terms, MR imaging starts by placing a subject in a relatively uniform, static magnetic field. The static magnetic field causes hydrogen nuclei spins to align with and cause a net magnetization in the general direction of the magnetic field. Radio frequency (RF) magnetic field pulses are then superimposed on the static magnetic field to flip some of the aligned spins, causing a net magnetization in a plane transverse to the static magnetic field that precesses about the field and thereby induces an RF response signal, called the MR echo or MR response signal. It is known that different tissues in the subject produce different MR response signals, and this property can be used to create contrast in an MR image. One or more RF receivers detect the duration, strength, frequency, and/or phase of the MR response signals, and such data are then processed to generate tomographic or three-dimensional images.
MR imaging can further provide a non-invasive means of quantitatively monitoring in vivo temperatures. This is particularly useful in MR-guided thermal therapy (e.g., MR-guided focused ultrasound (MRgFUS) treatment), where the temperature of a treatment area (e.g., a tumor to be destroyed by heat) should be continuously monitored in order to assess the progress of treatment and correct for local differences in heat conduction and energy absorption to avoid damage to tissues surrounding the treatment area. The monitoring (e.g., measurement and/or mapping) of temperature with MR imaging is generally referred to as MR thermometry or MR thermal imaging.
Among various methods available for MR thermometry, the proton resonance frequency (PRF) shift method is often the method of choice due to its excellent linearity with respect to temperature change, near-independence from tissue type, and temperature map acquisition with high spatial and temporal resolution. The PRF shift method is based on the phenomenon that the MR resonance frequency of protons in water molecules changes linearly with temperature (with a constant of proportionality that, advantageously, is relatively constant between tissue types). Since the frequency change with temperature is small, only −0.01 ppm/° C. for bulk water and approximately −0.0096 to −0.013 ppm/° C. in tissue, the PRF shift is typically detected with a phase-sensitive imaging method in which the imaging is performed twice: first to acquire a baseline PRF phase image prior to a temperature change and then to acquire a second phase image after the temperature change (hereinafter “treatment image”), thereby capturing a small phase change that is proportional to the change in temperature. A map of temperature changes may then be computed from the MR images by (i) determining, on a pixel-by-pixel basis, phase differences between the baseline image and the treatment image, and (ii) converting the phase differences into temperature differences based on the PRF temperature dependence while taking into account imaging parameters such as the strength of the static magnetic field and echo time (TE) (e.g., of a gradient-recalled echo).
The baseline phase can most easily be taken from the first image in a time series. This technique is, however, very sensitive to motion between the baseline images and images acquired during heating because motion causes misregistration between the images, resulting in erroneous baseline phase subtraction and inaccurate temperature estimates. If the treated organ's motion is cyclic, then an alternative to conventional baseline subtraction is acquisition and storage of multiple baseline images prior to therapy. During therapy, a well-registered baseline image can then be selected from the stored image library according to an image-similarity criterion, or based on additional information about the respective stages in the cycle of motion. Methods based on this concept are known as “multibaseline thermometry.” While more tolerant of motion than conventional baseline thermometry, this technique remains very sensitive to main-field shifts during therapy, which may result, for example, from respiration or bowel filling.
Another class of methods, collectively known as “referenceless thermometry,” is immune to both motion and main-field shifts. Referenceless thermometry utilizes no baseline images, instead deriving a reference phase image from the phase of the image portion corresponding to tissue surrounding a heated region. Operating under the assumption that the image phase directly surrounding a hot spot would otherwise extend smoothly into the region occupied by the hot spot, referenceless thermometry methods typically fit a set of smooth, low-order polynomial functions to the surrounding phase, or to a unit-magnitude complex image with the same phase. The polynomial is then extrapolated to locations within the hot spot, and the result used as a reference for subtraction. While referenceless methods are immune to motion, they are sensitive to rapid anatomical phase variations, since these cannot be accurately expressed as a weighted sum of smooth functions. Such variations commonly exist at organ edges. Furthermore, current referenceless methods require that the user know the location of the hot spot a priori, so that it can be masked out of the polynomial fit to avoid bias and temperature underestimation.
In light of the drawbacks of both multibaseline and referenceless PRF shift thermometry, alternative PRF methods that can compensate for motion and main-field shifts while being insensitive to anatomical boundaries are desirable.